X-ray systems are used for imaging for diagnostic examination and for interventional procedures, for example in cardiology, radiology and surgery. X-ray systems 16, as shown in FIG. 2, have an X-ray tube 18 and an X-ray detector 17, jointly arranged for example on a C-arm 19, a high voltage generator for generating the tube voltage, an imaging system 21 (often including at least one monitor 22), a system control unit 20 and a patient couch 23. Systems having two planes (2 C-arms) are also used in interventional radiology. Flat panel X-ray detectors are generally used as X-ray detectors in many fields of medical X-ray diagnostics and intervention, by way of example in radiography, interventional radiology, cardioangiography, but also the treatment for imaging within the framework of control and irradiation treatment planning or mammography.
Current flat panel X-ray detectors are generally integrating detectors and are predominantly based on scintillators whose light is converted into electrical charge in matrices of photodiodes. These are then read conventionally line-by-line by way of active control elements. FIG. 1 shows the principle construction of a currently used indirectly-converting flat panel X-ray detector, having a scintillator 10, an active reading matrix 11 made of amorphous silicon having a large number of pixels 12 (with photodiode 13 and switching element 14) and electronic control and reading device 15 (see for example M. Spahn, “Flat detectors and their clinical applications”, Eur Radiol. (2005), 15: 1934-1947). Depending on the radiation quality, the quantum efficiency for a scintillator made of CsJ with a layer thickness of for example 600 μm, depending on radiation quality, lies between about 50% and 80% (see for example M. Spahn, “Flat detectors and their clinical applications”, Eur Radiol (2005), 15: 1934-1947). The spatial frequency-dependent DQE(f) (“detective quantum efficiency”) is upwardly limited hereby and for typical pixel sizes of for example 150 m to 200 μm and for the spatial frequencies of interest to the applications of 1 to 2 lp/mm much lower. To enable new applications (for example Dual-Energy, Material-Separation), but to also further increase quantum efficiency, the potential of counting detectors or energy-discriminating counting detectors primarily based on directly-converting materials, such as CdTe or CdZTe (CZT) and contacted ASICs (application specific integrated circuit; for example embodiment in CMOS technology), is increasingly being examined.
FIG. 3 shows the basic construction of such counting detectors. X-ray radiation is converted in a direct converter 24 (for example CdTe or CZT) and the generated charge carrier pairs separated by an electrical field, which is generated by a shared top electrode 26 and a pixel electrode 25. In one of the pixel-like pixel electrodes 25 of the ASIC 27 the charge generates a charge pulse, the size of which corresponds to the energy of the X-ray quantum and which, if above a defined threshold value, is registered as a count event. The threshold value is used to distinguish an actual event from electronic noise or for example to also suppress k-fluorescence photons, in order to distinguish multiple counts. The ASIC 27, a corresponding section of the direct converter 24 and a coupling between direct converter 24 and ASIC 27 (in the case of directly-converting detectors for example by means of bump bonds 36) each form the detector module 35 having a large number of pixel elements 12. The ASIC 27 is arranged on a substrate 37 and connected to peripheral electronic devices 38. A detector module can also have one of more ASIC(s) and one or more parts of a direct converter, chosen as required in each case.
FIG. 4 shows the general diagram of a counting pixel element. The electrical charge passes the charge input 28, is collected in the pixel element and amplified there with the aid of a charge amplifier 29 and a feedback capacitor 40. The pulse shape can also be adjusted at the output in a shaper (filter) (not shown). An event is then counted in that a digital storage unit (counter) 33 is incremented by one if the output signal lies above an adjustable threshold value. This is established by way of a discriminator 31. The threshold value can in principle also be predetermined in a strictly analogue manner, but in general is applied across for example a DAC 32 (digital to analog converter) and can be variably adjusted in a certain region thereby. A reading can then be made by way of a control and reading unit 34. FIG. 5 shows a corresponding diagram for an entire array of counting pixel elements 12, for example 100×100 pixel elements each of, for example, 180 μm. Such an array is implemented with the aid of the ASIC. In this example it has a size of 1.8×1.8 cm2. For large-area detectors (for example 20×30 cm2) a plurality of detector modules 35 is combined (in this example 11×17 would produce roughly this area) and is connected by the shared peripheral electronic devices. TSV technology (through silicon via) for example is used for the connection between ASIC and peripheral electronic devices.
In the case of counting and energy-discriminating X-ray detectors two, three or more threshold values are introduced and the size of the charge pulse, corresponding to the predefined threshold values (discriminator thresholds), are classified in one or more of the digital storage unit(s) (counters). The X-ray quanta counted in a certain energy field may then be obtained by calculating the difference in the counter contents of two corresponding counters. The discriminators may be adjusted for example with the aid of DACs (digital-to-analog converter) for the whole detector module or pixel-by-pixel within given limits or ranges. The counter contents of the pixel elements are successively read module-by-module by a corresponding reading unit. This reading process requires a certain amount of time during which counting cannot continue without errors.
In a pulsed radiation mode of the X-ray system an X-ray radiation window and a reading window must be defined as in the case of integrating detectors. With an image frequency of for example 50 fps (frames per second; frame time for example 20 ms) and a reading time of 10 ms an X-ray window for a maximum pulse width of 10 ms remains (frame time=radiation time plus reading time). In this example (for example 10,000 pixels per detector module) a single pixel element is read at 1 μs (10,000×1 μs=10 ms). A design of this kind, which as in this example requires 1 μs reading time per pixel element, already reaches its limits at an image rate of 100 fps since the X-ray window then shrinks to 0 ms.
In the case of continuous radiation, the radiation process continues during the reading process and can consequently either not be counted (no use of the radiation dose) or, if this is not ruled out by the design, the count signals each correspond to slightly different periods of time, and this is something which should be avoided at all costs. Higher switching frequencies of for example 10 MHz can lower the reading time for the example detector module to 1 ms, but at 100 fps there is still a ratio of 1:9 between reading time and usable radiation time.